Total hip replacement prostheses have now been in use clinically for over 35 years. Most experience has been gained in the hip joint, where naturally, the prothesis is of the ball-head of the femoral stem and socket-type cup of the acetabular component. More recently reconstruction of the knee, elbow, shoulder and ankle have been attempted using prostheses of more varied mechanisms and design, although the basic concept for designing these prostheses remains unchanged.
One of the many universal requirements of implants, wherever they are used in the body, is the ability to form a suitably stable mechanical unit with the neighboring tissues. A loose (or unstable) implant may function less efficiently or cease functioning completely or it may induce an excessive tissue response. In either case it may cause t he patient discomfort and pain and in several situations a loose implant is deemed to have failed and has to be surgically removed. Total hip replacement prostheses have focused attention onto this stability question of the propensity for the load-transmitting prostheses to become over a period of time.
There are several ways in which the stability problem could be approached. One could consider a chemical approach in which adhesives were used to bond prostheses to tissues. Secondly, there is the mechanical approach and at least three different methods have been tried. One is now in routine clinical use and involves the interposition of a mass of acrylic "cement" into the space between prothesis and bone. There is no adhesive action as such, but rather a grouting effect which gives a mechanically stable unit and provides uniform load transfer.
Alternatively, one may choose an exceptionally inert material and prepare an accurate, well designed bed in the bone for the prothesis. The initial stability and the minimization of differentiation processes around the prothesis may allow new bone to grow right up to the prothesis surface to give a long-term stability. Using high density alumina ceramic appears to be a very bio-inert material and is a suitable candidate for these prostheses. However, as one might expect, the situation is not quite this simple and loosening can occur, even with very accurate bone-bed preparation. Surface macro-profiles on the prothesis may improve the stability but even here the limitations of this primary mechanical interlocking are shown.
In a third mechanical approach, the tissues do the work rather than the prothesis or a cement for this relies on tissue ingrowth into a micro-porous structure.
The important point is the extent of the fibrous tissue layer, if any, that forms between the material and the bone. It can be explained that it may need a biodegradable cement or temporary fixation to give initial stability while the necessary bone remodelling takes place to achieve a fibrous-tissue-free interface.
The development of porous implant was based on three key needs; biocompatibility through the use of ingredients (polytetrafluoroethylene and carbon fibers) that are stable in the body milieu and induce minimal biological disturbance through surface effects; biofunctionality through the selection of a modulus close to that of soft tissue and spongy bone (4 MPa); and maximum porosity (80 per cent void, 100.about.500 .mu.m diameter) and inter-pore connections (100.about.200 .mu.m) for tissue ingrowth consistent with appropriate bulk mechanical properties (ultimate tensile strength 1 MPa).
Referring to FIGS. 1 and 2, typical type of an orthopedic implant, particularly the total hip prothesis system, consists of a femoral prothesis 1 which has a long stem extending from a proximal portion with a ball-head portion to a distal portion and is inserted into a femur, and an acetabular socket-type cup 2 with an interlocking ring and insert which is implanted into an acetabulum. FIG. 3 shows an example of the total hip prothesis stem in which said femoral prothesis stem is interlocked into an acetabular cup (not shown).
It has been, for a long time, recognized that any type of implants (whether it is a dental implant or an orthopedic implant), should possess a biological compatibility against an implant-receiving surrounding hard and soft tissues. Accordingly, the materials for implants are limited to a certain type of materials, including a commercially pure titanium (CPT), T-6Al-4V alloy, Co--Cr or Co--Cr--Mo alloy, Fe--Cr--Mn--Ni--Mo--Nb--Si alloy, AISI Type 316L stainless steel, and some ceramic materials including pure alumina and the synthetic hydroxyapatite.
Similar type of materials used for the femoral prostheses can be found for the acetabular cups. One exception is that the ultra high molecular weight polyethylene (UHWMPE; average molecular weight is about more than 5,500,000) is a mainly selected material for the acetabular socket-type cup.
For these femoral prostheses and acetabular cups, several surface modifications have been proposed and used for enhancing the mechanical retention and promotion of the bony ingrowth.
In the ORTHOMET.RTM. system, very small size of commercially pure (CP) titanium beads are plasma-coated onto Ti-6Al-4V femoral stem to produce a porosity of about 50%. The CPT beads are coated on a proximal portion 3 in FIG. 3.
In the KYOCERA.RTM. system, CPT is coated onto a proximal portion of the Ti-6Al-4V stem material by the inert gas shielded arc spray to generate a coarse surface.
In the DePuy.RTM. system, very fine size of CPT beads are also coated on a proximal portion of the femoral stem to ensure a complete contact with the relatively weak cancellous bone structure.
In the development history of the bony ingrowth type artificial joint prostheses, there are two major design concepts available; namely, (1) a micro-fit type (or bone ongrowth type) in which grooves are formed on the surface of the implants, and (2) a micro-interlock type (or bone ingrowth type) in which the surface of implants is formed to be porous structure. In MULTILOCK.RTM. system, the mesh pad made of titanium fibers is bonded onto the proximal portion of the femoral prothesis through a solid state diffusion bonding technique.
In DePuy.RTM. system and STRAYKER.RTM. system, Co--Cr beads (approximately 150.about.300 .mu.m diameter) are sprayed or sintered to Co--Cr femoral stem material to generate an average pore size of 275 .mu.m.
In MATRIX HIP.RTM. system, the proximal portion of Co--Cr femoral stem was machined to create the macro- and micro-texturing for a reduced interfacial shear stress.
In ORTHOMET.RTM. system, polymethyl methacrylate (PMMA) resin studs are bonded as a pre-coating on an entire stem portion to enhance the bonding strength with the bone cement when the femoral stem is implanted into the cement mantle.
Moreover, in KYUCERA.RTM. system, synthetic hydroxyapatite is plasma-coated on the proximal portion of the stem to have an average particle size of 50 .mu.m. Alternatively, the proximal portion is subjected to a blasting to form a micro-textured structure.
Moving to the acetabular socket-type cups, in ORTHOMET.RTM. system, CPT beads are plasma-coated onto Co--Cr alloy to reduce the occurrence of the loosening.
In a similar manner, in STRYKER.RTM. system, CPT beads are bonded on the outer shell of the socket to enhance the mechanical retention.
In MULTILOCK.RTM. system, a plurality of PMMA studs are adhered on the outer surface of the acetabular head to promote a chemical bond with PMMA in the cement mantle.
In AcSysHA.RTM. system, synthetic hydroxyapatite is coated on an entire outer surface of the acetabular cup.
In the most outer shell of the acetabular head, a plurality of multi-hole, in-hole or cluster-hole are formed to promote the bony ingrowth activity.
Some of the aforementioned surface modification, particularly hydroxyapatite coating, are related to the mechanical comnatibility (which is the second compatibility required for the successful implant system). As discussed previously, the prostheses, particularly surface zone thereof, should respond to the load-transmitting function.
Suppose that an implant made of Ti-6Al-4V alloy is used. It is reported that strength and modulus of elasticity (MOE) of Ti-6Al-4V alloy are in ranges of 900.about.1200 MPa and 200.about.300 GPa, respectively. On the other hand, the strength and MOE of an implant-receiving bone has ranges of 100.about.200 MPa and 9.about.12 GPa, respectively. There are great differences in mechanical properties between these materials. With respect to the load-transmitting function, the strain field should be continuous. If the strain field in a metallic implant and bone system is not continuous, then the interface is not adhered and is easily debonded. However, even the strain field is continuous, the stress field has always a discontinuity due to the differences in MOE values between two dissimilar materials. If the discontinuous stress at an interface exceeds the interfacial bonding strength, then the bonding is failed.
Based on the above discussion, it is easily understood that the synthetic hydroxyapatite is plasma-coated onto the metallic implant, since the hydroxyapatite has not only an excellent biocompatibility, but it also exhibits the strength (400.about.800 MPa) and MOE (40.about.120 GPa) which are just between those values for Ti-6Al-4V and the implant-receiving bone structure. This is what is claimed as the second compatibility; namely, a mechanical compatibility to form a continuous gradient functional (in terms of mechanical property) structure.
Furthermore, the third compatibility should be included to the aforementioned currently accepted compatibilities; that is a morphological compatibility. In a scientific article published by this inventor (entitled "Fractal Dimension Analysis of Mandibular Bones: Toward a Morphological Compatibility of Implants" in Bio-Medical Materials and Engineering, Vol. 4 No. 5, pp.397/407, 1994), surface morphology of successful implants has a limitation of the average roughness (1.about.500 .mu.m) and average pore size (10.about.500 .mu.m), regardless of types of implant materials including metallic materials, ceramic materials and polymeric materials. If the pore size is smaller than 10 .mu.m, the surfaces will be more toxic to fibroblastic cells and has an adverse effect on cells due to their physical presence independent of any chemical toxic effects. If the pore size is larger than 500 .mu.m, the surface does not exhibit any sufficient structural integrity.
Moreover, as discussed previously, a mesh pad made of titanium beads is coated on the proximal portion of the femoral stem to promote the bony ingrowth. However, the opening space is not morphologically compatible to surrounding bone, particularly the cancellous bone. Any one of the beads (either made of Ti, hydroxyapatite, or PMMA) does not create sufficient opening space for the bony ingrowth and, then, is not morphologically compatible.
Successful implant system is relied not only on these three compatibilities, but also on cementation of the space gap between the prothesis and the vital bone.
With a certain type of the femoral prothesis, a ring-shaped cement riser made of PMMA is installed at a collar portion of the proximal part 4 in FIG. 3 and at a distal portion thereof 5 in FIG. 3. However, in the most femoral prostheses and acetabular cups, there is no pre-formed cement riser. Hence, during a surgical operation, cement mantle should be fabricated. Normally, the bone cement or bone paste consists of a mixture of PMMA resin and crushed bone autograft.
For a normal practice, the kneaded cement is used under a relative low viscosity. In order to keep the low viscous cement, the cement is cooled to prolong the wetting time until 90 seconds, which is advantageous, but the strength of the cement is slightly reduced. In order to compensate the reduced mechanical strength, the cement is mixed in vacuum to control the porosity.
Even with the aforementioned efforts, there are several important factors unsolved. These include (1) the heat generated during the polymerization of methyl methacrylate monomer, (2) the polymerization shrinkage, and (3) insufficient break strength and the modulus of elasticity.
Regarding the heat generated, the temperature rise is a function of the rate of polymer formation, or curing rate, rather than the rate of polymerization of a single polymer. Furthermore, the temperature rise is dependent of the rate of heat evolution and not necessarily upon the total heat evolved. Depending on the polymer formation rate, the temperature rise is reported in a range of 70 to 120.degree. C., which may irritate the surrounding hard and soft tissue and cause unexpected damage on them.
With regard to the polymerization shrinkage, it is reported that when methyl methacrylate monomer is polymerized, the density changes from 0.94 gm/cm.sup.3 to 1.19 gm/cm.sup.3. This change in density results in a volumetric shrinkage of 21 per cent, usually called the polymerization shrinkage. However, about one third in weight of the resin is monomer, then the volumetric shrinkage will be about 7% (that is approximately 2% of linear shrinkage). This amount of shrinkage is still too large and causes a dimensional accuracy and stability.
Regarding the mechanical properties, it is reported that the compressive strength of PMMA is about 50.about.70 MPa; while the bending strength is about 60.about.65 MPa. The modulus of elasticity in compression is 1.8.about.2.0 GPa; while MOE value under the bending mode is about 1.3.about.1.5 GPa. These are lower than those values for bone structure.